The present disclosure relates generally to a variable modulation scheme and more particularly to a variable x-ray power modulation scheme and a method for implementing the variable x-ray power modulation scheme.
In at least one known computed tomography (CT) imaging system configuration, an x-ray source projects a fan-shaped beam which is collimated to lie within an X-Y plane of a Cartesian coordinate system, wherein the X-Y plane is generally referred to as an xe2x80x9cimaging planexe2x80x9d. An array of radiation detectors, wherein each radiation detector includes a detector element, are within the CT system so as to received this fan-shaped beam. An object, such as a patient, is disposed within the imaging plane so as to be subjected to the x-ray beam wherein the x-ray beam passes through the object. As the x-ray beam passes through the object being imaged, the x-ray beam becomes attenuated before impinging upon the array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is responsive to the attenuation of the x-ray beam by the object, wherein each detector element produces a separate electrical signal responsive to the beam attenuation at the detector element location. These electrical signals are referred to as x-ray attenuation measurements.
In addition, the x-ray source and the detector array may be rotated, with a gantry within the imaging plane, around the object to be imaged so that the angle at which the x-ray beam intersects the object constantly changes. A group of x-ray attenuation measurements, i.e., projection data, from the detector array at one gantry angle is referred to as a xe2x80x9cviewxe2x80x9d. A xe2x80x9cscanxe2x80x9d of the object comprises a set of views made at different gantry angles during one revolution of the x-ray source and the detector array. In an axial scan, the projection data is processed so as to construct an image that corresponds to a two-dimensional slice taken through the object.
One method for reconstructing an image from a set of projection data is referred to as the xe2x80x9cfiltered back-projection techniquexe2x80x9d. This process converts the attenuation measurements from a scan into discrete integers, ranging from xe2x88x921024 to +3071, called xe2x80x9cCT numbersxe2x80x9d or xe2x80x9cHounsfield Unitsxe2x80x9d (HU). These HU""s are used to control the brightness of a corresponding pixel on a cathode ray tube or a computer screen display in a manner responsive to the attenuation measurements. For example, an attenuation measurement for air may convert into an integer value of xe2x88x921000 HU (corresponding to a dark pixel) and an attenuation measurement for very dense bone matter may convert into an integer value of +2000HU or more (corresponding to a bright pixel), whereas an attenuation measurement for water may convert into an integer value of OHU (corresponding to a gray pixel). This integer conversion, or xe2x80x9cscoringxe2x80x9d allows a physician or a technician to determine the approximate density of matter based on the intensity of the computer display.
Certain scanning parameters, such as x-ray tube, or emitter, current (xe2x80x9cmAxe2x80x9d), x-ray tube supply voltage (xe2x80x9ckVxe2x80x9d), slice thickness, scan time and helical pitch (for helical scans) are known to affect the x-ray power, which in turn affects image quality. In addition, the x-ray tube current typically directly relates to the patient x-ray dose. A higher x-ray tube current may, for example, improve the image quality but increase the dosage received by the patient. However, lower x-ray tube current levels are known to cause severe streaking artifacts in the image. This is typically caused by an insufficient number of photons passing through the patient and is known as x-ray photon starvation.
Although higher x-ray tube current levels result in lower noise images, the higher x-ray tube current levels subject patients to higher doses of x-ray energy. In conventional CT scanning practice, fixed mA protocols are used to scan a range of patients of various sizes and attenuation characteristics. As a result, the scans of smaller patients have less noise therein than the images of the larger patients. However, since a certain level of diagnostic image quality is required for larger patents, the smaller patients may therefore be receiving more doses than needed for acceptable diagnostic results when fixed mA protocols are used.
With regard to X-ray tube voltage, most CT scanners presently in use provide for several tube voltage stations (for example, 80 kV to 140 kV) that allow a technician and/or physician to adjust the x-ray tube voltage. However, voltage selection is mostly responsive to the preference of the physician, and thus typically lacks scientific guidance. For most body and head scans, some physicians tend to use a tube voltage of 120 kV, whereas others use 140 kV for head and pediatric scans where objects are relatively small. While on one hand higher tube voltage provides for better geometric dose efficiency for larger patients, lower tube voltage has been shown to provide for better contrasts for different types of lesions when the object is relatively small and, therefore, may provide for a better contrast to noise ratio (CNR). Unfortunately, these tradeoffs are not well established for medical practice and as such, the emitter tube voltage selection is generally fixed for a certain type of scan regardless of the patient size. Accordingly, it is desirable to be able to reduce the dose received by individual patients and to improve dose efficiency, while still maintaining acceptably small noise levels and good CNR.
The above discussed and other drawbacks and deficiencies are overcome or alleviated by a method for modulating the x-ray power of an imaging system so as to maintain a desired image noise in the imaging system. The method includes obtaining projection data and correcting the projection data responsive to beam hardening errors so as to create corrected projection data. In addition, the corrected projection data is processed so as to create a plurality of emitter current values responsive to an imaging method, and the emitter current values are applied to the imaging system responsive to an object to be imaged.
In another aspect, a method for determining an optimum emitter tube voltage for an imaging system includes characterizing the imaging system so as to determine a system water-equivalent path length responsive to a relative noise increase. An object water-equivalent path length is then determined and compared with the system water-equivalent path length so as to create a comparison result, allowing for the recommendation of the optimum emitter tube voltage responsive to the comparison result.
In another aspect, a system for modulating the emitter current of an imaging system so as to maintain a desired image noise in the imaging system includes a gantry having an x-ray source and a radiation detector array. The gantry defines a patient cavity, wherein the x-ray source and the radiation detector array are rotatingly associated with the gantry so as to be separated by the patient cavity. A patient support structure is movingly associated with the gantry so as to allow communication with the patient cavity. A processing device obtains projection data and corrects the projection data responsive to beam hardening errors so as to create corrected projection data. The processing device processes the corrected projection data so as to create a plurality of emitter current values responsive to an imaging method and applies the emitter current values to the imaging system responsive to an object to be imaged.
In still another aspect, a system for determining an optimum emitter tube voltage for an imaging system includes a gantry having an x-ray source and a radiation detector array. The gantry defines a patient cavity, wherein the x-ray source and the radiation detector array are rotatingly associated with the gantry so as to be separated by the patient cavity. A patient support structure is movingly associated with the gantry so as to allow communication with the patient cavity. A processing device characterizes the imaging system so as to determine a system water-equivalent path length responsive to a relative noise increase. The processing device further determines an object water-equivalent path length and compares the object water-equivalent path length with the system water-equivalent path length so as to create a comparison result, and recommends the optimum emitter tube voltage responsive to the comparison result.
The above discussed and other features and advantages of the present invention will be appreciated and understood by those skilled in the art from the following detailed description and drawings.